The present invention relates to the art of active electro-magnetic shielding. It finds particular application in conjunction with self-shielded magnetic gradient coils used in magnetic resonance imaging (MRI), and will be described with particular reference thereto. However, it is to be appreciated that the present invention is also amenable to other like applications where high quality efficient active shielding of magnetic and/or electric fields is desired.
In MRI systems, gradient coil assemblies are commonly pulsed with electrical current pulses to produce magnetic gradients across a main magnetic field in the vicinity of an imaging region. Commonly, the magnetic field gradients interact with external metallic structures such as the magnet cold shields, the magnet dewar, and the like. This interaction tends to generate eddy currents in the affected structures. These eddy currents, in turn, generate eddy magnetic fields which have a deleterious effect on the temporal and spatial quality of the magnetic field in the vicinity of the imaging region and, hence, in the resultant image quality.
The eddy current problem is often addressed by placing an active shielding coil between the primary gradient coil and the affected structure. The shielding coils are designed to substantially zero or cancel the residual magnetic field external to itself thereby preventing the formation of eddy currents in potentially vulnerable structures.
Previously methods for production for magnetic gradients in MRI systems consisted of winding discrete coils in a bunched or distributed fashion on an electrically insulating hollow cylindrical former and driving the coils with a current source of limited voltage. Conventional bunched coil designs include the Maxwell and the Modified Maxwell Pair for z-gradient production (i.e. axial gradient production), and the Golay or Modified Golay (multi-arc) Saddle Coils for x and/or y-gradient production (i.e. transverse gradient production). Typically, these methods consisted of iteratively placing coil loops or arcs on the cylindrical formers until the desired gradient strength, gradient uniformity, and inductance (related to stored energy) were achieved. These previous designs were generally developed in a "forward approach" whereby a set of initial coil positions were defined (i.e., the initial coil distribution), the fields and the inductance/energy calculated, and if not within particular design parameters, the coil positions would be shifted (statistically or otherwise) and results re-evaluated. This iterative procedure continued until a suitable design was obtained.
More recent methods of generating magnetic fields in magnetic resonance imaging systems utilize an "inverse approach." In the "inverse approach" method, the gradient magnetic field is forced to match predetermined values at specified spatial locations inside the imaging volume and a continuous current density is calculated which is capable of producing such a field. The "inverse approach" method assumes that the primary gradient coil has finite dimensions while those of the secondary or shield coil are left unrestricted (infinite). After the generation of continuous current distributions for both the primary and the shield or secondary coils, an apodization algorithm is performed on the continuous current density of the shield coil in order to restrain it to desirable dimensions. Following the modification of the shielding coil's continuous current, the Stream Function technique is employed in order to obtain discrete current patterns for both coils. Application of the Biot-Savart law to the discrete current pattern ensures that the discretization procedure was proper. This approach creates generally more energy efficient gradient coil assemblies with higher gradient strengths and faster slew rates as compared to the "forward approach" method.
One particular prior art approach is discussed in U.S. Pat. No. 4,794,338 to Roemer, et al. The approach of designing a shielded gradient coil assembly introduced by Roemer, et al. is based on the "forward approach" method. The outcome is a shielded gradient coil assembly with a moderate to low efficiency rate in terms of gradient strength and slew rate. Further, there is no precondition to the method for controlling eddy current effects inside the imaging region.
Another particular prior art gradient coil assembly is described in U.S. Pat. No. 5,296,810 to Morich. Morich describes a cylindrically shaped shielded gradient coil assembly for MRI applications. Morich uses the "inverse approach" method of designing gradient coil assemblies where the primary coil has a finite length while the length of the shielding coil is considered infinite. This configuration generates coils with high gradient strengths and slew rates while at the same time reduces the eddy current effects when the length of the shield coil is substantially longer (20% or more) than the length of the primary coil. In order to restrain the current of the shielding coil within desired dimensional boundaries, apodization techniques (e.g., guassian apodization) are employed. In this manner, the overall length of the shielding coil is approximately 20% longer than the total length of the primary gradient coil. Since the shielding coil was modeled as having an infinite length, in order to confine the length of the shielding coil to finite dimensions, current truncation is employed. The current apodization process towards the ends of the shielding coil cause disturbances in the shielding field and ultimately result in unwanted eddy current effects inside the imaging region. These effects become more deleterious as the length of the shielding coil approaches the length of the primary coil, or as both coil dimensions approach the imaging region.
In another particular prior art shielded gradient coil assembly described in U.S. Pat. No. 5,132,618 to Sugimoto, the design is based on the "inverse approach" method. In this design, both the primary and the shielded coil lengths were assumed infinite and continuous current densities for both the primary and the secondary coil are modeled based on this assumption. In order to restrain the current densities on both the primary and secondary coils, truncation is again employed. Although the outcome of this method is similar to that of the Morich patent discussed earlier, the additional truncation of the primary coil's current in this case introduces increased levels of eddy current effects inside the imaging region.
For interventional procedures and like applications where patient access is desirable, it is advantageous to design the gradient shielding coil such that its dimensions do not exceed those of the primary gradient coil. In this manner, patient access is maximized and the feeling of openness reduces patient claustrophobia. However, in general, the previous methods and prior art discussed above suffer the drawback that as the shielding coil length approaches that of the primary coil, increased levels of eddy current effects within the imaging region deteriorate image quality. Conversely, when sufficient shielding is achieved, the dimensions of the shielding coil are sufficiently larger than those of the primary coil that the level of patient access is encumbered and an increased level of patient claustrophobia is experienced.
Additionally, prior methods commonly determine the current density for the primary coil first, and then determine the appropriate current density for the shielding or secondary coil which will produce the zeroing of the residual magnetic field outside itself. These approaches, in varying degrees, often fail to appreciate the cooperative and interactive nature of the two-coil system. Consequently, they achieve generally less favorable results.
The present invention contemplates a new and improved technique for actively shield magnetic and/or electric fields which overcomes the above-referenced problems and others.